PET scanners are well known in the field of medical physics. These scanners produce images of the body by detecting radiation emitted from radioactive substances injected into the body. Each scanner is made up of radiation detectors, typically called scintillators, which are arranged in a ring configuration around a movable patient table. A typical arrangement with a detector ring 10 and a patient table 12 is shown in FIG. 1. Each scintillator comprises a crystal and has an associated partner located opposite it on the ring. Many known cameras use Bi4Ge3O12 (BGO) as a scintillation detector, as taught in U.S. Pat. No. 4,843,245 and EP 0,437,051 B. Each scintillator is connected to a photomultiplier tube, which is in turn connected to read-out electronics.
During a scan the patient is positioned on the movable table in the centre of the ring of detectors. The patient is injected with a radioactive substance, which is tagged with a β+ radioactive atom that has a short decay time, for example carbon-11, fluorine-18, oxygen-15 or nitrogen-13. During decay of the nuclei of the radioisotopes, positrons are given off. When a positron is emitted and meets an electron, the collision produces two gamma rays that have the same energy, 511 keV, but travel in opposite directions. By detecting coincidentally the gamma rays generated using scintillators that are diametrically opposed on the ring, the trajectory on which the disintegration occurred can be detected. The scintillator crystals convert the gamma rays to photons of light that are transmitted to the photomultiplier tubes, which convert and amplify the photons to electrical signals. These electrical signals are then processed by a computer to generate three dimensional images of the body over the region of interest (e.g. brain, breast, liver).
An advantage of PET scanning is the ability to determine accurately radionuclide localization and to quantify physiological processes in the body. This can be done because of the emission from the patient's body of two gamma photons that travel in opposite directions. Another advantage is that PET scanners use biological compounds similar or identical to those found in the human body, such as carbon, nitrogen, and oxygen. This means that the PET radionuclides can be substituted directly into biological substances used by the body. In addition, it means that PET tracers do not merely mimic biological pathways as do agents for some other scanners, instead PET tracers actually follow true physiological and metabolic processes. This is advantageous. In contrast, other nuclear medicine imaging techniques require compounds labelled with radioactive nuclides not commonly found in the body. These modified compounds only approximate the true distribution in the human body.
Because of the many advantages inherent in PET scanners, there is a drive to improve their performance, thereby to increase the accuracy of the scanned images and so assist clinicians. To this end, work is presently being done by many groups to improve the characteristics of such scanners.
The most important characteristics of a PET camera are its spatial resolution and sensitivity. Conventional PET cameras can provide spatial resolutions in the range of 4–6 mm at full width half maximum (FWHM) of the emission spectrum. Better spatial resolution requires a large number of scintillation detectors with reduced size, and as a consequence, a large number of photodetectors and associated read-out electronics. This, however, increases the cost. At the same time, new demands on human PET instrumentation, for example on precise brain imaging, require spatial resolution to be better than 2 mm.
The combined formula for a reconstructed image resolution for a PET scanner can be expressed as follows:Γ=1.254√{square root over ((d/2)2+(0.0022D)2+r2+b2)}{square root over ((d/2)2+(0.0022D)2+r2+b2)}Here Γ is the reconstructed image resolution in mm FWHM, d is detector size, D is the detector array diameter, which is typically 600–800 mm for a whole body PET scanner and 250–300 mm for a brain PET (NB including D takes into account photon non-collinearity from positron decay), r is the effective positron range (from 0.5 mm for 18F to 4.5 mm for 82Rb), and b is an additional factor, which is derived from a hit point identification scheme (Anger Logic or “true” position sensitive photo detector, i.e: analog ratios among many photomultiplier signals). Assuming that b is zero for a position sensitive photodetector, it is possible to achieve (with 18F) Γ=1 mm resolution for brain a PET at d=1 mm.
In addition to limitations on the spatial resolution of known PET scanners, there is also a geometrical limitation on the spatial resolution at the end of a PET scanner field of view. This is the so-called radial elongation distortion, which occurs when gamma trajectories cross several scintillation detectors.
Assuming that the above equation is accurate, in order to satisfy the growing demand on the spatial resolution and sensitivity of PET cameras, it will be appreciated that the camera has to be made of thin detectors with high stopping power. In practice, however, the stopping power being limited by the density of the material the detector needs to have a certain length, typically one to three centimetres which reduces the spatial resolution at the edge of the field of view. This is a disadvantage. To overcome this problem and avoid degradation of the spatial resolution, it is necessary to use a detector with depth of interaction (DOI) determination capability, i.e. the ability to determine the interaction coordinate along the detector cell. The most convenient way to do this is to use a multi-layer detector, in which the layers are made of material with different scintillation properties. Because the layers have different characteristics, when a gamma ray is detected it is possible to identify the layer that was hit and so determine more accurately the interaction point.
Many multi-layer detectors are known. For example, U.S. Pat. No. 4,843,245 describes a multi-layer scintillator that uses adjacent BGO and GSO (Gd2SiO5) crystals. EP 0,219,648 teaches the use of a three layer scintillator that has an inner layer of BaF2, a middle layer of GSO and an outer layer of BGO. WO 99/24848 also teaches the use of a multi-layer detector and in particular a “phoswich” detector, in which different detector layers are made of different scintillators with different decay times. The phoswich described in WO 99/24848 has two layers, one each of BGO and Lu2SiO5:Ce (LSO).
Another known multi-layer detector uses a combination of LSO and GSO. In this case, hit layer determination is carried out using pulse shape discrimination. This can be done because of the large difference in decay time constants of the LSO and GSO layers. Unfortunately, the photoelectric absorption coefficient of GSO is much less than that of LSO. This means that the stopping power of the GSO is limited, which introduces a degree of uncertainty into the determination of the hit layer.
In yet another known PET, the scintillation detectors are made of layers of “fast” and “slow” LSO scintillators, grown with different cerium concentrations.
As with the LSO and GSO detectors, pulse shape discrimination is used to determine the hit layer. A disadvantage of this particular device is, however, that the difference in decay time constants of “fast” and “slow” LSO is only about 10% (4–5 nanoseconds at mean value of 40 ns). Hence, there can be difficulty in determining the hit layer with any certainty.